1. Field of the Invention
The present invention generally concerns nuclear magnetic resonance tomography (MRT) in the field of medicine for examination of patients. The present invention in particular concerns a method and a magnetic resonance tomography apparatus to implement the method of the type making use of the two-point Dixon method.
2. Description of the Prior Art
Magnetic resonance tomography is a slice image method for medical diagnostics that is primarily characterized by a high contrast resolution capability. Due to the excellent ability to show the soft tissue, magnetic resonance tomography has developed into a method that is in many cases superior to x-ray computed tomography. Magnetic resonance tomography today is based on the application of spin echo and gradient echo sequences that enable an excellent image quality given measurement times on the order of minutes.
Obtaining fat images and water images of a patient represents a challenge in magnetic resonance imaging. Due to the influence of the chemical shift (see below), artifacts that must be corrected arise at the boundary layers between fat and water, but the fat signal represents valuable information (for example fat content in the liver) which should be maximized or optimized.
The property that the resonance frequency shifts slightly in proportion to the field strength depending on the type of chemical bond in which a signal-emitting nucleus participates, is known as chemical shift. Due to their concentration in the human body, hydrogen atoms in free water and in fat primarily contribute in the image. Their relative resonance frequency difference is approximately 3 ppm (parts per million). In the use of spin echo and gradient echo sequences, this leads to a modulation of the signal intensity depending on the echo time TE.
In the original publication by W. T. Dixon, a method was presented that achieves a separation of fat images and water images with two echoes (gradient or spin echoes). This is described briefly in the following.
The magnetization vector of the water protons Mw and the magnetization vector of the fat protons Mf point in the same direction immediately after radiation of a radio-frequency excitation pulse (typically with a flip angle of 90° in the case of a spin echo sequence, typically with much smaller flip angle given a gradient echo sequence). However, this state does not persist since the water protons in the homogeneous magnetic field precess 3 to 4 ppm more quickly than the fat protons. In a laboratory system (FIG. 2) it is seen how the magnetization of the water protons and that of the fat protons disperse with time. This difference amounts to approximately 50 Hz at 0.35 T. As shown in FIG. 3, the total magnetization MT is the vector sum of water magnetization and fat magnetization. FIG. 3 is based on a reference system that rotates with the frequency of the water protons.
FIG. 4 shows that the total magnetization MT initially (when the water magnetization and the fat magnetization point in the same direction) exhibits a maximum but soon traverses a minimum when the water magnetization and the fat magnetization are anti-parallel.
The first minimum occurs when
  t  =            1              2        ⁢                  (                                    v              W                        -                          v              F                                )                      =    a  wherein t is thereby the time, vF is the fat proton frequency and vw is the water proton frequency. The time a is of great importance since the acquisition of an imaging sequence at time t=a delivers an image in which the brightness of the pixel depends on the difference between fat magnetization and water magnetization. An acquisition at t=2a (an echo can not yet be acquired at t=0 since this must first form during t=2a; the fat-water magnetization is anti-parallel at t=3a), thus when fat magnetization and water magnetization are aligned in parallel, yields an image in which the sum of fat magnetization and water magnetization is shown.
The sum and the difference of the two images are now of decisive importance: the sum yields a water image, the difference yields a fat image. It is noted that both images still exhibit an additional, system-dependent phase at the point in time of the measurement. The correction of this phase is necessary but need not be explained in detail within the scope of the discussion herein.
The method described has a significant disadvantage: it does not take into account that both echoes (gradient echo or spin echo) are negatively affected by different decay processes (relaxation processes), namely that the gradient echo is typically affected by the different location-dependent relaxation time of the transversal magnetization that is characterized by T*2, and the spin echo is typically affected by the plain T2 decay (thus without taking into account the local B0 field inhomogeneities). Actual existing inhomogeneities due to relatively strongly T*2-dependent or T2-dependent relaxation processes lead to the situation that the components cannot be separated without doubt. In the further proceedings, it is essentially T*2-sensitive gradient echo techniques that are considered, without limitation of generality.
The solution to this problem is presently avidly researched. All present solution approaches are based on the measurement of additional echoes, typically more than three (up to eleven echoes can be measured). This has the disadvantage of a significant lengthening of the measurement time that is particularly unacceptable for most clinical applications. Moreover, the image resolution is also drastically reduced.